Nuclear magnetic resonance is sometimes used to obtain information about the chemical composition of matter, e.g., as a radio frequency spectrometer.
Nuclear magnetic resonance is also used as a technique for obtaining images of patients' body tissues. In this context, it is sometimes referred to as magnetic resonance imaging. In either case, it may be used as a noninvasive measurement technique for medical purposes in the treatment of patients.
Magnetic resonance imaging relies on the principle that hydrogen atoms, when subjected to a magnetic field, line up like so many soldiers. If a radio frequency signal is aimed at these atoms, it changes the alignment of their nuclei. When the radio waves are turned off, the nuclei realign themselves, transmitting a small electric signal. And since the body is primarily composed of hydrogen atoms, an image can be generated from the returning pulses, showing tissue and bone marrow.
Nuclear magnetic resonance ("NMR") imaging offers advantages over x-ray imaging. The magnetic fields used in NMR imaging are not harmful like x-rays. Moreover, x-ray imaging only has one degree of freedom. X-ray imaging relies upon the electron density, or relative absorption of x-rays between various tissue groups. It is sometimes impossible to detect, for example, brain tumors using x-ray imaging because the diseased tissue has virtually the same x-ray absorption as the surrounding normal tissue.
In contrast to x-rays, NMR imaging may be said to have three degrees of freedom: proton density and the relaxation times T.sub.1 and T.sub.2.
These additional degrees of freedom permitted by NMR imaging may be advantageously used to obtain information and to distinguish between tissues, which would not be possible with other techniques.
In some cases, a brain tumor may not be visible in an x-ray image. But NMR imaging may be able to distinguish between such tumor tissue and normal brain tissue, even where x-ray imaging would fail to detect any difference between the two tissue groups. In addition, NMR imaging can be used to look at heart tissue and to detect damaged tissue due to the buildup of lactic acid levels, where x-ray imaging might not be able to detect such tissue.
Much of the hydrogen present in body tissue is found in the form of water molecules. Each molecule of water contains two hydrogen atoms. Bones contain little water, and therefore do not appear in NMR images. As a result, NMR imaging enables doctors to see tissue that is surrounded by bone. NMR imaging is an especially effective tool for imaging the spinal cord. Previously, doctors who wanted to look at the spinal cord using x-ray imaging had to inject it with an x-ray contrast agent, typically using a painful and risky procedure.
In NMR measurement, it is desirable to increase the signal-to-noise ratio of the system. This has been an impetus in the development of higher field and higher frequency apparatus. The signal-to-noise ratio may be increased by increasing the frequency under certain conditions.
For purposes of providing a context for the present invention, it is helpful to start with a general description of NMR measurement techniques. In an NMR application, an essentially static magnetic field is applied to a patient's tissue. A sample coil is placed in close proximity to the patient's tissue which is desired to be imaged. A pulse of radio frequency electric current is then applied to the sample coil at a particular frequency, in order to effect a maximum NMR decay signal.
An NMR scanner surrounds the body with powerful electromagnets. Typically supercooled by liquid helium, these electromagnets create a strong magnetic field.
This field has a profound effect on protons, the nuclei of hydrogen atoms. Spinning like tops, the protons normally point in random directions. But inside the NMR scanner's magnetic field, they generally align themselves in the direction of the magnetic field's poles. Even in alignment, however, they wobble, or precess, at a specific rate, or frequency. The stronger the magnetic field, the greater the frequency (f+).
When the scanner excites these protons with a radio pulse timed to the same frequency as their wobbling, it knocks them out of alignment. Within milliseconds they spiral back into place, "singing out" with a faint radio signal of their own.
A computer translates these faint signals into an image of the area scanned. The image reveals varying densities of hydrogen atoms and their interaction with surrounding tissues in a cross section of the body. Since hydrogen reflects water content, doctors can use the image to make distinctions between tissues. Hydrogen is commonly selected as the basis for NMR image scanning because of its prominent magnetic qualities. Other elements may also be employed, such as sodium or phosphorous, whose altered properties could provide early warning signs of strokes or heart attacks. It may be possible to tag cloned antibodies with a detectable element as a tool to study such disorders as diabetes, allergies, infertility, and cancer.
An NMR imaging system may be described generally in connection with FIG. 12. FIG. 12 is an exploded perspective drawing illustrating a static field magnetic coil 20, and X, Y, and Z gradient coils 61, 62 and 63, respectively. In an NMR measurement system, the gradient coils 61, 62 and 63 are placed within the static field magnetic coil 20, and a patient 64 is placed in the cylindrical area inside of the coils.
To make an image, a control computer 31, (see FIG. 13), establishes a grid of tiny boxes, or voxels, in three dimensions, X, Y and Z. First the magnetic field is varied in the Z direction, from head to toe, to define a plane of interest where the body 64 will be scanned. Within this plane protons wobble at a given frequency, F. Radio frequency ("RF") coils 22 then emit a pulse at precisely the same frequency to topple these protons.
Before the protons can realign themselves, other coils 62 briefly vary the magnetic strength of the plane in the Y direction. This causes protons to wobble at different rates from the top of the plane to the bottom. Detecting these differences over many pulse-and-response cycles, the computer 31 locates voxels in the Y direction.
Coils 61 then vary the magnetic field from left to right in the X direction, causing protons to "sing" at different frequencies as they realign themselves. A similar process is performed using the Z gradient coils 63. Having located each voxel in the X, Y and Z directions, the computer 31 assigns each voxel a spot on a video screen. The spot's brightness is determined by the number of protons within the voxel and the magnetic properties of the tissue. Together the dots form a readable image.
An RF pulse is generated by an RF transmitter 27a, and then fed to the RF coil 22 through an RF switch 24. After the pulse has been transmitted, the RF switch 24 switches the RF coil 22 to the RF receiver 27b. The RF coil 22 then acts as a receiving antenna to pick up the signals emitted by protons when the protons "sing" as they realign themselves.
Further description of magnetic resonance imaging appears on pages 14-23 of the January 1987 issue of National Geographic, the entirety of which is incorporated herein by reference. Further background information appears in U.S. Pat. No. 4,599,565, issued to Hoenninger, III, et al., entitled "Method and Apparatus for Rapid NMR Imaging Using Multi-Dimensional Reconstruction Techniques", the entire disclosure of which is incorporated by reference. Also, the entire disclosure of U.S. Pat. Nos. 4,634,979, 4,583,044 and 4,585,992 are all incorporated herein by reference.
In a more technical description of nuclear magnetic resonance, NMR may be described as a technique by which anatomical images and biochemical information can be obtained noninvasively from a patient's body. Unlike conventional x-ray CT-scan technology, no harmful ionizing radiation is used. Instead, harmless static magnetic fields and short pulses of radio frequency energy is used to induce desired species of atomic nuclei in the body to a state of resonance. This nuclear resonance generates a resonant magnetic field which is received from body tissues during the time intervals between RF pulses. This received signal termed free induction decay ("FID") is then detected, sampled, and digitally processed into the desired presentation format, whether it be anatomical images, or chemical spectra.
Other than being more safe, NMR imaging and NMR chemical spectroscopy are superior to x-ray CT technology in several important ways. NMR spectroscopy can provide in-vivo metabolic information, which x-rays cannot do. NMR imaging, more commonly referred to as magnetic resonance imaging ("MRI") is capable of much more resolving power than can be had with x-ray images. X-ray imaging relies upon the relative electron densities of different tissues to contrast them. MRI generally relies on the hydrogen nucleus or relative proton densities to contrast between tissues. Although other atomic nuclei such as fluorine or phosphorous may be used for imaging as they are in spectroscopy, hydrogen protons are most often used because of their abundance in the body, hence strong FID signal. In addition to the relative abundance of protons in tissues, T.sub.1 and T.sub.2 relaxation times may be used to further contrast different tissues. T.sub.1 is a property of the molecular structure of a tissue, and T.sub.2 is a property of the number of differences in molecular structures of a tissue. T.sub.1 is the time required for resonant protons to relax back to an nonresonant state during the receive period following the excitatory RF pulse. Generally, the more rigid tissues like bone and muscle have shorter T.sub.l 's than the more fluid tissues like blood and lymph. T.sub.2 relaxation is due to a relative dephasing of the poles of precessing protons over time, thereby causing a decay in the signal level received. So with NMR's chemical information available plus three degrees of freedom in image contrasting versus one degree of freedom for x-ray, tissues may be resolved and contrasted with NMR where x-ray techniques fail. Small tumors, or tumors with similar densities to their host tissue cannot be detected with x-rays, but can be detected with NMR. Regions of tissue death or disfunction as in a stroke in the brain, or an infarct in the heart can often only be detected with NMR. Soft tissues in general can be resolved using NMR to a much higher degree than can be done with x-rays.
The part of the NMR system that is used to transmit the RF pulse, and to receive the FID signal from the patient, is an RF tuned antenna commonly referred to as a coil. The coil is a resistive-inductive-capacitive ("RLC") circuit tuned to the static magnetic field dependent resonant frequency of the desired atomic nuclei. The traditional coil structure is usually an inductive loop of wire whose terminals are in electrical parallel with a variable capacitor. Capacitance is varied to tune the coil to the desired nuclear resonant frequency for maximum energy transfer, much like tuning a radio or television to a desired frequency for maximum signal reception. For a given static magnetic field, each nuclei has its own characteristic resonant frequency by the relation: EQU .omega.=.gamma.B.sub.0
where .omega. is the resonant frequency, .gamma. is a nuclei specific constant, and B.sub.0 is the field strength of the static field. A coil is placed near the surface of a patient such that the axis of the coil is perpendicular to the axis of the magnet within which the coil and patient are positioned. For the direction of the RF pulse coupled to the terminals of the coil, a B.sub.1 resonant magnetic field is stored in space about the inductive loop of the coil. The homogeneous portion of this B.sub.1 field in the center of the coil is perpendicular to the static B.sub.0 field of the magnet. The nuclei specified by the resonant frequency of the coil align as would tiny bar magnets with the resonant B.sub.1 field of the coil. After the RF pulse is terminated, the nuclei relax or realign with the B.sub.0 field of the magnet. The spinning or precessing net alignment of the poles of the nuclei population is termed the magnetization vector. This magnetization vector is induced by the RF pulse to the coil during transmission, and then in turn induces an RF current in the coil during the FID signal reception period between pulses as the vector realigns with the B.sub.0 field.
This FID signal is very low, (typically measured in microvolts), especially for nuclei less abundant and less sensitive to NMR than protons. The signal-to-noise ratio ("SNR") is therefore critical. An increase in the SNR translates directly into more NMR data. More body chemistry may be seen. Higher resolution, higher contrast images may be obtained. Of all NMR system components, the coil placement and design has the most impact on the SNR. A coil is electrically loaded when the coil is placed near the patient surface. Loaded coils have been shown theoretically and empirically to adhere to the following proportionality: ##EQU1## where .eta. is the degree to which the coil is loaded, M is the magnitude of the magnetization vector, Q is defined as the energy stored in the field about the coil divided by the power dissipated in the patient and in the coil, and .omega. is the resonant frequency of the nuclei and the tuned coil. As we saw earlier, .omega. is directly proportional to B.sub.0, the magnetic field strength. V is the volume of the coil.
To improve the SNR of a coil, we wish to increase Q, .omega., and V by coil design. .eta. is dependent on the coil proximity to the patient. M is dependent on the magnetically coupled nuclei. In the past, for traditional single loop, or series loop coils, one could not increase Q, .omega., and V simultaneously, but had to sacrifice one variable for an improvement in another. To increase V, we increase the inductance of the coil, thereby decreasing Q and .omega.. If the field strength of the magnet used in the NMR system is increased, .omega. increases. But to increase .omega. for a conventional coil, V is lost. Most conventional series type coil designs used in the past were not capable of achieving the higher resonant frequencies required by higher B.sub.0 field magnets. Finally, to increase Q, the size of the coil V had to be decreased.
With the present invention, all three parameters Q, .omega., and V can be simultaneously increased to improve the SNR of a surface applied coil. This invention involves a substantially planar array of two or more electrically resonant ("RLC") loops connected in electrical parallel with respect to each other and their transmit/receive source. These loops are connected such that all are excited with equal phase and equal RF energy. These loops can be individually switched, or attenuated to produce a desired B.sub.1 magnetic field distribution. This net field distribution is the superposition of the fields from each loop.
Connecting N loops together in a parallel array permits five advances in surface coil technology: (1) an increase in V by a factor of up to N; (2) an increase in .omega. by a factor of up to N; (3) an increase in Q by a factor of up to N; (4) an overall increase in SNR; and, (5) an improvement in magnetic field spatial localization capabilities.
The strong magnetic fields and large volumes used in state-of-the-art NMR in-vivo applications require the advances in coil technology listed above.
In nearly every NMR application, the sample coil is made to resonate at the frequency used. Normally, the sample coil is tuned to resonance using an external capacitor. For a given coil, there is a frequency above which it is difficult to attain resonance because the capacitance cannot be made small enough. The resonant frequency is inversely related to the square root of the inductance multiplied times the capacitance. Thus, if one attempts to use a larger coil in order to obtain a higher field strength to improve signal-to-noise ratio, the highest frequency at which that coil may be effectively used will be severely limited. These limitations encountered in old coil designs constitute a significant problem which is addressed by the present invention.
Attempts have been made in the past to overcome deficiencies in surface coils used in nuclear magnetic resonance measurements. One such attempt involved the use of surface coils referred to as loop-gap resonators. An example of such an attempt is described in an article published in Medical Physics by James S. Hyde, et al., entitled "Planar-Pair Local Coils For High-Resolution Magnetic Resonance Imaging, Particularly Of The Temporomandibular Joint." A two-loop, one-gap resonator is shown. The loops formed inductance, and the gap provided capacitance.
Currents in the loop, however, were forced by the potential across the gap to be in opposite directions. Thus, the magnetic fields produced within each loop were out of phase. This caused a shadow to be formed in the image between the two loops because the magnetic fields generated by the two loops destructively interfered with each other. These so-called loop-gap resonators were stated by the authors to be useful for imaging nominally symmetrical structures on either side of the body mid line in those cases where the intervening tissue is not of interest. While the authors claim such loop-gap resonators may be useful to image orbits, breasts and/or kidneys, the limited applicability of such a device is a severe disadvantage. It is desirable to have an improved surface coil configuration which may be used to image an increased area of patient tissue, including all intervening tissue. In most applications, the intervening tissue normally is of interest.
In the Hyde et al. article, the authors state that with two loops the inductance doubles and the apparent resistance (due to eddy currents) doubles, resulting in an unchanged Q. The authors also state that because the effective resistance is doubled, the signal-to-noise ratio is reduced. Thus, the so-called loop-gap resonators were not even claimed to be effective in improving the signal-to-noise ratio.
In an article published in the Journal of Maqnetic Resonance by Stephen B. W. Roeder, et al., entitled "NMR Coils With Segments In Parallel To Achieve Higher Frequencies Or Larger Sample Volumes", the authors describe an apparent attempt to connect coils in parallel. However, the coils were arranged axially. In this article, the authors state that dividing a coil into a number of parallel loops or subcoils should result in a coil with a Q reduced, at most, by the number of subcoils.
Contrary to the teachings of the Roeder et al. article, the present invention allows the Q of planar surface coils, electrically connected in parallel, to result in a Q which may be increased by a maximum amount which may approach as a limit the number of subcoils which are connected in parallel. The inductance may be reduced, thereby simultaneously allowing an increase in the resonant frequency. The area of the imaging field which is provided by the parallel planar surface coils is also increased.
In the prior art, there has been no disclosure of how to increase the resonant frequency, increase the Q, and increase the volume V covered by a surface coil, all at the same time. One feature of the present invention is that it allows all three of these factors to be increased simultaneously.